By way of an electrochemical analysis method, substances can be determined both qualitatively and quantitatively on account of specific physical properties using the electric current. Electrochemical analysis methods in which electrode reactions play a part are of particular importance. These are classified into two groups depending on whether the excitation signal (current, voltage or potential) is kept constant. By way of example, potentiometry, chronopotentiometry, coulometry, amperometry, dhronoamperometry and chronocoulometry are techniques in which the excitation signal is kept constant. In voltammetric and polarographic methods, the excitation signal is varied.
Together with optical methods such electrochemical analysis methods for the analytical determination of chemical and biochemical substances are characterized by a high sensitivity and also a high selectivity. Whereas, however, complicated, expensive and sensitive optical and optoelectronic apparatuses are necessary in the case of optical analysis methods, electrochemical analysis methods manage with comparatively simple electrode devices.
An advantage of electrochemical analysis methods is the direct presence of the measurement result as an electrical signal. The latter, after analog-to-digital conversion, can be processed further directly by a computer, preferably by a personal computer.
Electrochemical analysis methods are suitable for the quantitative and qualitative measurement of substance concentrations in an electrolyte solution. Every substance has an oxidation voltage and reduction voltage, respectively, that are characteristic of the substance. It is possible to distinguish between different substances on the basis of these voltages. Furthermore, the concentration of a substance present can be deduced on account of the electric current that flowed during a reaction.
An electrochemical experiment requires at least two electrodes which are connected to the substance to be analyzed in electrolytic solution (working electrode, counterelectrode). However, it is also possible to use a plurality of working electrodes in parallel. Usually a reference electrode is additionally used for exact supervision of the electrolyte potential. This system having at least three electrodes is connected to a potentiostat, which makes it possible to regulate the potential at the working electrode and measures the electric current flowing through the working electrode.
In the case of voltammetry, a variable voltage is applied to the working electrode and the current flowing during an oxidation or reduction is measured. In the special case of cyclovoltammetry, a specific voltage range is repeatedly swept over in such a way that the constituents of the electrolyte are repeatedly successively oxidized and reduced.
In the case of chronoamperometry, a defined voltage is applied discontinuously to the working electrode and the current that flows is recorded over time. This measurement method permits the analysis of a specific substance by targeted oxidation or reduction of said substance. The current that flowed is a measure of the quantity of substance converted per unit time and permits conclusions to be drawn with regard to the concentration of the substance and with regard to the diffusion constant.
Chronocoulometry corresponds to chronoamperometry in terms of the electrical boundary conditions. In contrast thereto, however, the total electrical charge that flowed is recorded rather than the electric current that flowed.
In the configuration as sensors, electrode devices can be used in various electrochemical analysis methods. All that is crucial is that substances that can be evaluated electrochemically are generated during the sensor event. In the case of sensors for the detection of biomolecules, use is made of a marking method, by way of example, which provides electrochemically active substances in the case of a sensor event.
Miniaturized electrochemical electrode systems for the analysis of chemical and biochemical substances are known in the prior art, see for example [1], [2], [3], [4] and [5]. The electrodes of such arrays can be individually contact-connected at the edge of the substrate and be operated by means of a potentiostat. In order, moreover, to realize electrode arrays which have for example 100 or more electrodes, switching functions on the substrate which multiplex the electrodes onto common connecting lines are advantageous. If the substrate is a semiconductor material, such as silicon, the required switches may be realized by MOS transistors, as described in [6]. On account of the parallelization that can thereby be achieved in the tests, the analysis time is significantly shortened and it also becomes possible to carry out complex analyses.
Reference [7] describes the so-called EDDA method (Electrically Detected Displacement Assay method), from the company Friz BIOCHEM™.
Reference [8] describes an electrode system for detecting molecules or molecular complexes. The assembly contains at least three electrodes, at least one working electrode, one counterelectrode and one reference electrode being present. The reference electrode is arranged in such a way that it is adjacent at least to partial regions of the two further electrodes. Furthermore, an electrode assembly described in [8] contains a plurality of operational amplifiers integrated into a substrate and serving for evaluating a sensor signal, which operational amplifiers supply the measurement signals to external evaluation units.
Reference [9] describes a biosensor for detecting macromolecular biopolymers. An electrode assembly having at least one unit for immobilizing macromolecular biopolymers is formed in the biosensor. In this case, the at least one unit is provided with catcher molecules. Furthermore, the at least one unit is set up in such a way that a medium which is to be examined and which contains biopolymers to be detected can be brought into contact with the immobilization unit. Biopolymers contained in the sample to be examined can be bound to the catcher molecules, whereby complexes composed of biopolymers and catcher molecules are formed. A first electrical measurement is carried out in order to detect the biopolymers with the aid of the biosensor. Afterward, the complexes are separated and a second electrical measurement is carried out. The biopolymers are detected by way of an alteration of the value of the electrical measurement that is brought about by the separation of the complexes.
Reference [10] describes a DNA sensor having an electrode assembly with an interdigital structure. In this case, the interdigital structure has additional reaction areas for attachment of thiols. During the operation of the sensor described in [10], so-called markers are applied to the areas covered by thiols.
Reference [11] describes an electrode assembly for detecting macromolecular biopolymers which has at least one unit for immobilizing biopolymers. The at least one unit for immobilizing macromolecular biopolymers is provided with catcher molecules, in which case the catcher molecules can bind macromolecular biopolymers, on the one hand, and on the other hand have a marking that can generate a detectable signal.
Reference [12] describes a circuit for the integration of an electric current (current integrator), which circuit has an operational amplifier, inter alia.
Reference [13] describes a circuit for the rectification of an AC voltage. In this case, a capacitor is charged during a positive half-cycle of the AC voltage by way of a forward-biased diode. During the negative half-cycle, the diode is turned off, and the voltage present at the capacitor remains approximately constant.
From the standpoint of miniaturization, signal integrity and also measurement sensitivity, the active microarrays known in the prior art constitute very good electrochemical analysis systems [2]. Such electrochemical sensor arrays that operate according to voltammetric, (chrono)amperometric and (chrono)coulometric methods are manufactured in accordance with CMOS technology and equipped with electrodes made of a noble metal (e.g. gold) which are accessible on the chip surface.
Active sensor arrays have been realized in DNA sensor chips, for example, in which, on the basis of redox cycling, DNA molecules are detected at surfaces electronically by the detection of electrical charge carriers generated by means of redox-active substances. Redox cycling constitutes a special case of an amperometric method (oxidation/reduction voltages constant, measurements of the electrode current).
A typical redox cycling sensor arrangement has two gold electrodes formed on a substrate. Single-stranded DNA catcher molecules immobilized via the so-called gold-sulfur coupling, for example, and having a predetermined sequence are immobilized on each electrode. The complementary single-stranded DNA target molecules, which are thus capable of hybridization, possibly present in the analysis solution have a marking. By way of the marking, given the presence of suitable additional molecules, a cycle of oxidation and reduction of constituents of the additional molecules is initiated, which, under interaction with the electrodes, leads to the formation of reduced or oxidized molecules. The cycle of oxidations and reductions leads to a circulating electric current that enables a detection of the DNA target molecules.
Both in the case of said redox cycling assembly and in the case of the electrochemical analysis methods mentioned above, a counterelectrode is always required. Whereas in the case of the redox cycling sensors, however, only a relatively small DC current has to be conducted away at the electrodes, in the case of most of the electrochemical analysis methods specified above it is necessary for a comparatively high surge current to be able to be supplied by the counterelectrode. For this reason, the area of the counterelectrode has to be significantly greater than that of the active working electrodes.
Depending on the concrete analysis method, a surface area approximately 10 times greater with respect to the sum of the surface areas of the individual working electrodes regularly has to be demanded for the counterelectrode. This is necessary because if the counterelectrode has an excessively small area, the voltage present at it during an experiment for providing the required charge carriers can assume extremely high values. If such high values are assumed, chemical conversions, e.g. of the electrode material, that proceed in uncontrolled fashion may be the consequence, which typically take place with the formation of gases.
If the surface area of the counterelectrode is large enough, it is able to stabilize the electrolyte potential for the most part by way of the double-layer capacitance. Electrochemical conversions take place only with comparatively low current densities.
Since an active silicon chip as a substrate, e.g. for a DNA sensor, is comparatively expensive, generally a highest possible packing density of the individual sensors in the array is striven for. Owing to the packing density of the sensors and thus of the electrodes, under certain circumstances it is not possible to realize a counterelectrode in the region of the sensor array. The counterelectrode may then be embodied as an external electrode which is arranged in the sample volume and electrically connected to the sensor chip. The driving of this electrode may be performed by a potentiostat. This procedure is disadvantageous, however, owing to comparatively long leads and the more complicated mechanical construction. If the associated disadvantages are to be avoided, the only solution offered by the prior art is to realize the counterelectrode in the periphery of the array, but this requires additional (expensive) chip area.
One important type of sensor, particularly in the case of all-electronic DNA sensor chips, is based on so-called redox cycling. Principles of redox cycling are described in [3], [4]. Redox cycling involves detecting macromolecular biopolymers at surfaces electronically by detection of electric currents caused by way of redox-active markings.
FIG. 1A, FIG. 1B show a redox cycling sensor assembly 100 in accordance with the prior art.
The redox cycling sensor assembly 100 has two gold working electrodes 101, 102 formed on a substrate 103. DNA catcher molecules 104 having a predetermined sequence are immobilized on each working electrode 101, 102. The immobilization is effected in accordance with the so-called gold-sulfur coupling, by way of example. Furthermore, an analyte 105 to be examined is introduced into the redox cycling sensor assembly 100. The analyte 105 may be for example an electrolytic solution comprising various DNA molecules.
If the analyte 105 contains first DNA single strands 106 having a sequence which is not complementary to the sequence of the DNA catcher molecules 104, then the first DNA single strands 106 do not hybridize with the DNA catcher molecules 104 (see FIG. 1A). This case is referred to as a “mismatch”.
If, by contrast, the analyte 105 contains second DNA single strands 107 having a sequence which is complementary to the sequence of the DNA catcher molecules 104, then the second DNA single strands 107 hybridize with the DNA catcher molecules 104. This case is referred to as a “match”. To put it another way, a DNA single strand 104 having a predetermined sequence is in each case only able to hybridize selectively with a very specific DNA single strand, namely with the DNA single strand having a sequence that is complementary to the respective catcher molecule.
As shown in FIG. 1B, the second DNA single strands 107 to be detected contain a redox-active marking 108. After the hybridization of the second DNA single strands 107 to be detected with the DNA catcher molecules 104, by way of the redox-active marking 108 (e.g. an enzyme label such as e.g. an alkaline phosphatase), given the presence of suitable additional molecules 109 (for example para-aminophenyl phosphate, p-APP), a cycle of oxidations and reductions of constituents of the additional molecules 109 is initiated, which, under interaction with the gold electrodes 101, 102, leads to the formation of reduced molecules 110 (e.g. para-aminophenol) and oxidized molecules 111 (e.g. quinone imine). The cycle of oxidations and reductions leads to a circulating electric current that enables a detection of the second DNA single strands 107.
Consequently, in the redox cycling method, in the case of a binding event between a particle to be detected and a catcher molecule by means of an enzyme label (e.g. an alkaline phosphatase), a redox-active species is produced by means of para-aminophenyl phosphate (p-APP), contained in an electrolyte for example being converted into para-aminophenol. Since new redox-active species are constantly generated, this leads to a rise in the electric current between the two electrodes.
An oxidizing electrical potential is required at the first working electrode 101, which may also be referred to as generator electrode. A reducing electrical potential is required at the second working electrode 102, which may also be referred to as collector electrode.
FIG. 2 shows an interdigital electrode arrangement 200 known from the prior art, which assembly has two interdigitated working electrodes, namely a generator electrode 201 and a collector electrode 202. A reference electrode 203 and a counterelectrode 204 are furthermore shown. The electrodes 201 to 204 are formed on a substrate 205. An electrolytic analyte (not shown) may be applied to the interdigital electrode arrangement 200, said analyte being coupled to the electrodes 201 to 204.
The electrical potential of the electrolytic analyte is provided, by way of the reference electrode 203, at an inverting input of a comparator 206 and compared by the latter with a desired electrical potential at the noninverting input of the comparator 206. In the case where the electrical potential of the reference electrode 203 deviates from the desired potential, the counterelectrode 204 is driven via an output of the comparator 206 in such a way that said counterelectrode subsequently supplies electrical charge carriers as required in order to maintain the desired electrical potential of the electrolyte. The reference electrode 203 together with the comparator 207 clearly forms a potentiostat device. The electrical potentials at the working electrodes 201, 202 are set relative to the reference voltage. First and second ammeters 207, 208 are used to detect electric sensor currents of the generator electrode 201 and of the collector electrode 202, respectively, which contain items of information about a sensor event that has possibly taken place.
The prior art furthermore discloses a sensor array in which a plurality of interdigital electrode assemblies 200 are connected up to one another in matrix-type fashion, for example. In said array, components 203, 204, 207, 208 may be provided jointly for a plurality of sensor zones.
Circuit architectures for biosensors are known which serve for the sensitive detection of biomolecules by way of electrochemical conversions in electrolyte solutions. In this case, in particular the hybridization of two complementary oligonucleotides is detected by means of the presence of known electrochemical markings (e.g. a Ferrocen marking). The electrical measurement of the electrochemical markings present on the surface of the sensor electrode is effected by means of an abrupt change in the electrode voltage. In this case, the markings are oxidized or reduced in a targeted manner. The value of the current that occurs in the process is a measure of the quantity of the marking present on the sensor surface. In this case, the current surge or the quantity of charge is composed of electrical charges which flow into the double-layer capacitance at the electrode surface and the charges which emerge from oxidation and reduction processes.
In accordance with the prior art, an absolute value of an electrical charge quantity that occurs during a voltage surge is measured at one sensor electrode. What is disadvantageous in this case is that large quantities of such electrical charge carriers which do not contribute to the measurement signal and thus impair the noise margin are also detected in an undesirable manner.
Reference [14] describes an all-electronic DNA sensor array chip that uses a redox cycling method, the individual sensors of the array in each case having an electrode assembly with an interdigital structure and also a potentiostat circuit. The DNA sensors has 128 sensor positions, an analog-to-digital conversion of the respective sensor signal being effected in each individual pixel. The chip is realized using a CMOS process.